Magnetic resonance imaging (MRI) takes advantage of the magnetic properties exhibited by many atomic nuclei and the presence of these atomic nuclei within a sample tissue to create detailed images of internal structures. In the absence of an external magnetic field, the spins of magnetic nuclei are oriented randomly. However, when placed within a strong, static magnetic field (B0), the spinning nuclei adopt one of two orientations—they either align with the B0 field or against the B0 field. The first orientation, in which a particular nucleus is aligned with the B0 field, is slightly lower in energy than the second orientation, in which a nucleus is aligned against the B0 field. An MRI apparatus typically employs a primary electromagnet to generate the strong, static B0 field. By exciting the nuclei in the first orientation with a RF magnetic field (B1) at a frequency that corresponds to the energy differential between the first orientation and the second orientation, the nuclei in the first orientation absorb energy and “flip” to the second orientation. After excitation, the nuclei undergo a transition back to the first orientation. During this transition, the nuclei emit the absorbed energy as RF signals. These signals are captured by a RF receiver and are used to compile the image.
The difference in energy between the first and second orientations is based on the type of atom and the strength of the B0 field. It is desirable to perform MRI imaging utilizing a higher strength static magnetic field B0 because higher strength B0 fields lead to increased signal to noise (SNR) ratio. However, increased static field strengths also lead to certain problems. For example, because the energy differential between the first and second orientations increases with increasing static field strength, so too does the RF energy required to induce the transition from the first orientation to the second orientation. As such, the frequency of the RF signal to generate the B1 field is increased. By way of example, RF energy of approximately 64 MHz is required to bring a 1H nucleus into resonance at a static field strength of 1.5 Tesla whereas RF energy of approximately 300 MHz is required at a field strength of 7 Tesla.
MRI systems utilize radio frequency transmit coils to produce the B1 excitation field. These coils are designed in varying shapes, sizes, and configurations depending on the particular application. For example, a surface coil is contained in a single plane and is placed in proximity to a region to be imaged. In contrast, a volume coil, such as a birdcage coil or a solenoid coil, surrounds a volume to be imaged. At the higher frequencies required to induce the transition from the first orientation to the second orientation at higher static field strengths, interactions between the RF field and the tissue being imaged result in non-uniform B1 fields in the tissue. This B1 field inhomogeneity reduces the quality of images obtained by the MRI system. A well-known solution to provide a more uniform B1 field throughout a volume of tissue being imaged utilizes multiple RF transmit coils (i.e., a transmit coil array) positioned about the volume rather than a single RF transmit coil. By optimizing the relative magnitudes and phases of the currents on the multiple coils, the B1 field can be made more homogeneous.
While the positioning of coils in a transmit array can produce a more uniform B1 field throughout a volume of tissue being imaged, transmit arrays also lead to certain problems. In order to produce a uniform B1 field, it is necessary to control the current flowing through each coil of the array. However, because the RF pulse is generally defined by a voltage level input to an amplifier, the unique loading of each of the coils may lead to different currents on the coils. For example, the properties of the tissue that is located in proximity to a particular coil will affect the impedance of that coil. Therefore, coils in different positions will have different impedances based on their proximity to tissues having different properties. To further complicate matters, the coils in a transmit array may be inductively coupled such that a change in current flowing through one coil induces a voltage on another coil. The mutual inductive coupling of coils in a transmit array makes it even more difficult to control the current delivered to a coil in the array.
One proposed solution to control currents on each coil of a transmit array utilizes a separate isolation power amplifier to supply each coil. Isolation power amplifiers differ from conventional power amplifiers in that they effectively present a blocking impedance to the coil. This essentially removes the effects of mutual inductive coupling, allowing each coil to be programmed separately. While this method may allow each coil to be independently programmed, there are certain drawbacks to this approach. First, because MRI transmit coils may operate at “non-standard” frequencies and high power ratings, isolation power amplifiers with the required characteristics may not be common items and may therefore be quite expensive. Moreover, a separate isolation power amplifier is required for each coil in the transmit array. Second, while the isolation power amplifier solution may remove the effects of mutual inductive coupling, each coil in a transmit array will still exhibit different loading characteristics based on the properties of the coil and its positioning. Thus, each coil in a transmit array must be separately tuned to provide equal currents to each of the coils. While this is a workable solution, it is desirable to supply each coil in a transmit array from a common amplifier and to tune the system as a whole. There is therefore a need in the art for an MRI system in which a single amplifier can deliver uniform currents to each coil in a transmit array.